The present invention relates to the art of nuclear medical imaging. It finds particular application in conjunction with rotating one-dimensional (1D) slat-collimated gamma cameras and single photon emission computed tomography (SPECT), and will be described with particular reference thereto. However, it is to be appreciated that the present invention is also amenable to other like applications and other diagnostic imaging modes, such as positron emission tomography (PET).
Nuclear imaging employs a source of radioactivity to image the anatomy of a subject. Typically, a radiopharmaceutical is injected into the patient. Radiopharmaceutical compounds contain a radioisotope that undergoes gamma-ray decay at a predictable rate and characteristic energy. Various scanning techniques exist in which the emitted xcex3-rays are detected. Based on information such as detected position and energy, the radiopharmaceutical distribution is located in the body and a representation of some feature of the subject, such as an organ, abnormality, etc., is reconstructed.
In a traditional Anger-type camera, the detector includes a scintillation crystal that is viewed by an array of photomultiplier tubes. The heads have collimators disposed between the crystal and the subject to limit the trajectory along which radiation can be received. Typically, the collimators are thick lead plates with an array of apertures or bores. Radiation traveling in a trajectory through one of the bores strikes the crystal; whereas radiation traveling in other trajectories hits the collimator and is absorbed. In this manner, each scintillation defines a ray, typically perpendicular to the face of the crystal, although magnifying and minifying collimators are also known. The thicker the collimator, the more accurately the ray trajectory is defined, but count efficiency or sensitivity is reduced since more radiation is absorbed in the collimator without reaching the detector.
Rather than using a single, large scintillator and photomultiplier tubes, others have proposed using an array of small scintillators, each associated with a photodiode or other photosensitive device which senses a scintillation in each individual scintillation crystal. Other types of individual solid-state detectors have also been suggested.
To improve the amount of radiation that reaches the detector, it has been proposed to use collimator sheets in a single direction across a row of detectors such that detected radiation defines a plane instead of a ray. The detectors are rotated to collect the planes at many angles. For three-dimensional images, the detector was positioned at a plurality of locations around the subject and the rotating data collection process repeated.
Solid state radiation detectors, such as cadmium-zinc-telluride (CZT) detectors, cadmium-telluride detectors, and the like, are also known, which utilize the photoelectric effect to detect radiation. That is, received radiation photons liberate electrons from their orbits around atoms of the target material. A high bias voltage is applied across the detector material to aid the photoelectric phenomenon and electron propagation. The electrons are detected as an electrical signal. Although very good performance can generally be expected from room-temperature CZT, sometimes a pixel is defective, for example, due to crystal impurities, crystal boundaries, electrical contacts, and other reasons.
In a conventional two-dimensional array, a dead pixel can hardly be tolerated and, techniques are known to avoid xe2x80x9cholesxe2x80x9d in the image, such as substituting the value of an adjacent pixel, substituting an xe2x80x9caveragexe2x80x9d value of pixels neighboring the dead pixel, etc. However, such techniques degrade spatial resolution and sensitivity. In the case of pixels having poor energy resolution, their presence, too, degrades performance of the two-dimensional array, although an ill-behaved pixel is generally more tolerable than no pixel at all.
The process of selecting and testing CZT crystals for two-dimensional arrays adds a significant cost to an already expensive technology and might, in practice, lead to a substantial relaxation of the performance criteria.
The present invention provides a new and improved method and apparatus that overcome the above referenced problems and others.
In accordance with one aspect of the present invention, a nuclear imaging apparatus comprises a radiation detector including a plurality of rows of detector elements, which generate an output pulse in response to each detected radiation event. A rotor rotates the radiation detector and a plurality of summing circuits, each connected with one of the detector element rows, generate a sum of the output pulses therefrom during a sampling period. A correction circuit adjusts the sums with correction factors, each row having a preselected correction factors. A reconstruction processor reconstructs an image representation from the adjusted sums and rotational position of the detector information corresponding to each sampling period.
In accordance with another aspect of the present invention, a nuclear imaging apparatus includes a radiation detector comprising an array of solid state detector elements responsive to incident gamma radiation by emitting a current spike. A pixel correction processor detects defective detector elements in the array and a flood correction circuit corrects detected radiation events based on sensitivity differences between a plurality of groupings of detector elements in the array. A reconstruction processor reconstructs an image representation from the corrected radiation events.
In a further aspect, a method of diagnostic imaging includes exposing a solid state radiation detector to a known radiation source. The radiation detector comprises a two-dimensional array of detector elements or pixels generating a detectable signal responsive to incident gamma radiation, and each detector element comprising a distinct channel. Radiation events are detected at each detector element and defective and nondefective pixels are identified and detector element correction values are calculated to normalize the energy spectrum of each nondefective pixel. An energy window is defined based on the normalized spectra to distinguish those photons having the energy characteristic of the radiation source. Also, a weighting factor for each row of pixels is calculated to scaling each row to a nominal value when the detector is exposed to a known radiation source. A radioactive isotope is injected into a subject located in an imaging region and the detector array is rotated while detecting radiation events indicative of nuclear decay. Multiple planar projections are generated of an examination region at a plurality of angular orientations, wherein radiation events from the defective pixels is either not recorded or discarded. The detector array is moved around a longitudinal axis of the subject and the steps of rotating and detecting are repeated. The detected radiation events are collected by row and corrected with the weighting factors to generate corrected data. An image representation of the subject is reconstructed using the corrected data.
In yet another aspect, a method of calibrating a nuclear imaging device includes exposing a solid state radiation detector to a known radiation source. The radiation detector comprises a two-dimensional array of detector elements or pixels generating a detectable signal responsive to incident gamma radiation, and each detector element comprising a distinct channel. Radiation events are detected at each detector element and defective and nondefective pixels are identified and detector element correction values are calculated to normalize the energy spectrum of each nondefective pixel. An energy window is defined based on the normalized spectra to distinguish those photons having the energy characteristic of the radiation source. Also, a weighting factor for each row of pixels is calculated to scaling each row to a nominal value when the detector is exposed to a known radiation source.
Another advantage of the present invention is that a gamma camera with improved performance is provided, especially in terms of energy resolution.
Another advantage is that the cost of the array is substantially decreased by allowing the use of detector crystals that would ordinarily be rejected due to the presence of one or more bad pixels.
Another advantage of the present invention is that dead and defective pixels can be turned off to improve energy resolution without any loss of spatial resolution and only a minimal effect on sensitivity.
Still further benefits and advantages of the present invention will become apparent to those skilled in the art upon a reading and understanding of the preferred embodiments.